Optical endoscopes are shown to be devices of increasing interest for the investigation of the human body. State of the art endoscopes can be classified into two categories: widefield endoscopes based on fiber bundles or GRIN lenses and scanning based systems using single mode optical fibers. In both cases, these systems are limited either by their size, flexibility, imaging resolution or their weak light collection. Most endoscopic systems are adapted from commonly used free space microscopy techniques where the image contrast can be generated in different ways.
An image can be, for example, obtained by collecting light scattered by a sample where the changes in the optical refraction index gives structural information. This linear scattering imaging can be improved by applying a specific illumination (as in bright-field or dark-field imaging). Biological structures can also naturally emit light by fluorescence when they absorb a probe light. By collecting the so-called auto-fluorescence signal, an image can be constructed. Artificial fluorescent markers can also be added to the sample for targeting a specific structure and obtain functional information about that structure. The fluorescence signal, emitted by these markers after excitation, can be collected to form what is called an optical fluorescence image.
Optical fluorescence imaging relies on a sample (dyed or not) emitting a fluorescence signal after light excitation. An image is obtained by collecting the fluorescence emission, which is usually at a longer wavelength than the excitation, either through a wide field microscope or a scanning optical microscope.
In wide field microscopy, the fluorescence image is directly collected and formed through the optical system. The endoscopic equivalent relies on the use of either a short (mm scale) small diameter GRIN lens as a microscope objective or fiber bundles, which consists of fibers arranged in an array. In the latter, the resolution is limited by the inter-core spacing.
Scanning optical microscopes, however, are based on producing an image point by point by scanning a diffraction limited focus spot. For each scanned position at the sample, the optical information, which can be either a linear signal (scattering, single fluorescence) or a non-linear signal, is collected to form an image. The detected light resulting from one illuminated volume element represents one pixel in the resulting image. The beam is scanned across the sample in two dimensions or in three dimensions (axially). For depth selectivity or sectioning, a pinhole can be added to form a confocal arrangement. In a confocal scanning microscope, the illumination/excitation beam first passes through an aperture and then is focused on the sample. Scattered and reflected laser light as well as any fluorescent light from the illuminated spot is then re-collected by the objective lens. The detection apparatus has a pinhole that obstructs the light that is not coming from the focal point. The out-of-focus light is rejected leading to a sharp-er image and giving the possibility to perform optical sectioning by acquiring images at various depths. A confocal arrangement can be made with GRIN-lens endoscopes but not with a fiber bundle.
Other imaging techniques based on non-linear effects are so-called two-photon imaging (or multi-photon imaging) and Raman imaging. In two-photon imaging, excitation is based on the effect that two photons of comparably lower energy than needed for one photon excitation, can also excite a fluorophore in one quantum event. Each photon carries approximately half the energy necessary to excite the molecule. An excitation results in the subsequent emission of a fluorescence photon. Since the probability of simultaneous absorption of two photons is extremely low, a concentrated flux of photons is necessary, a condition experimentally realized at the focus of a high numerical objective and using femtosecond pulsed laser sources. Two-photon imaging systems typically require fluorophores tagged to a specimen of interest in order to obtain strong two-photon efficiency. GRIN-lens endoscopes and fiber bundles have been demonstrated to provide dispersion compensation to maintain a short pulse duration at the sample.
Raman imaging uses the Raman effect which is an inelastic scattering effect in which a probe photon (from a probe beam) interacts with the vibrational levels of the probed molecules. The resulting scattered photon is energy-shifted by an amount equal to the energy of the vibrational level involved. Raman scattering is weak (typ. 1 ppm) and thus a high number of photons per volume is needed to produce a Raman shifted photon. This condition is experimentally realized at the focal spot of a lens. Continuous wave laser beams are typically used. An enhancement of the Raman signal is realized by a nano-patterned surface (metal) on which the probed molecules reside. The electric field at specific spots on the nano-patterned surface is enhanced by a plasmonic effect (electron resonance effect with the light frequency). The so-called surface enhanced Raman scattering (SERS) is proportional to the incident optical electric field to the power four at resonance. Both two-photon and Raman techniques are based on point-by-point measurements and thus a scanning system is required to form an image as in a scanning optical microscope.
In microscopy, the lateral spatial resolution d of a sample is limited by the wavelength of light λ, and the numerical aperture NA of the microscope objective via the Abbe relation: d=λ/2NA. A better resolution can be achieved if one uses a so-called “super-resolution”imaging microscopy technique. Two major techniques are used that both rely on a patterned illumination: STimulated Emission Depletion microscopy (STED) which is a scanning based method and Structured Illumination Microscopy (SIM), which is a wide-field method.
STED (stimulated emission depletion) microscopy makes use of non-linear de-excitation of fluorophores to overcome the Abbe diffraction limit and was proposed by Hell and co-workers (U.S. Pat. Nos. 5,731,588, 7,064,824, 7,430,045). With this technique, a structure is tagged with a substance such as a fluorophore that can be in either of two states having each a specific optical property. The state of the substance can be toggled between the first and second state and vice-versa by means of a switch-over optical signal. The light induced toggling between the two states is non-linear with light intensity. A first excitation beam is focused by a high numerical aperture objective lens on the sample tagged with the fluorophores to bring the latter to an excited state. A second beam, red-shifted to the first optical beam is focused by the same objective to form a doughnut beam at its focus. The red-shifted beam toggles the fluorophores to a second state by a stimulated depletion effect. The doughnut beam has zero intensity only at the center and thus fluorophores, located within an area smaller than the diffraction limit around the center, remain in the first state. Fluorophores in the first state emit a fluorescence radiation that is collected by the same high numerical objective and separated by color filters from the first and second excitation/de-excitation beams respectively. To form an image, the Gaussian and doughnut spot size are scanned together, e.g., by a system of rotating mirrors placed before the high numerical objective. The lateral spatial resolution d in STED is dependent on the light intensity of the de-excitation beam ISTED: d=1/NA*sqrt(1+ISTED/Isat).
SIM microscopy is, unlike STED microscopy, a wide-field technique that can improve the resolution of a fluorescence light microscope by at least a factor of two (U.S. Pat. No. 6,239,909; US 2012/0026311). SIM uses a grid to create several interference patterns on the sample. The illumination pattern interacts with the fluorescent probes in the sample to generate interference patterns known as moiré fringes that include high-resolution information that is normally inaccessible. These moiré patterns are superimposed upon each other to form a single image. This can be done by using widefield microscopy and placing a fine mesh grating in the light path before excitation. In Fourier optics, the resolution limit is defined by the optical transfer function, which is the normalized Fourier transform of the point-spread function. When two frequencies are mixed (the multiplication of two signal), moiré fringes are created. When moiré patterns are combined, information outside of the conventionally observable space becomes observable.
Other super resolution microscopy techniques are based on a stochastic illumination, as in PALM (Photo Activated Localization Microscopy) and STORM (Stochastic Optical Reconstruction Microscopy). They utilize sequential activation and time-resolved localization of photoswitchable fluorophores to create high-resolution images. During imaging, only an optically resolvable subset of fluorophores is activated to a fluorescent state at any given moment, such that the position of each fluorophore can be determined with high precision by finding the centroid position of the single-molecule images of particular fluorophore. The fluorophore is subsequently deactivated, and another subset is activated and imaged. Iteration of this process allows numerous fluorophores to be localized and a super-resolution image to be constructed from the image data.
The microscopy techniques mentioned above rely on free-space optical components such as high numerical aperture microscope objectives and can be adapted, in a more or less straightforward way, to an endoscopic device. Scanning-based endoscopic devices usually rely on the use of a single-mode fiber for the focused excitation and a second fiber, usually multimode, for light collection. In the case of the wide field technique, GRIN rigid lenses are used or bundles of large number of single-mode fibers are used, even if they have an inherent resolution limitation coming from the inter core spacing.
With respect to super-solution techniques, STED, PALM and STORM have never been implemented in an endoscopic device but SIM has already been implemented using a fiber bundle.
In Raman imaging, fiber probes exist in which the probe beam is transmitted to the sample under test by a single mode fiber (U.S. Pat. No. 5,112,127). Because of the tight light confinement in the single mode core of the fiber, a strong Raman signal is generated in the fiber itself. To mitigate this effect, a larger core fiber is used. However, this is achieved at the expense of resolution. This Raman signal needs to be optically removed from the main beam as this Raman signal (elastically scattered by the sample) can mask the Raman signal generated by the sample. Because of this Raman induced signal in the fiber, a small lens is placed at the distal end of the fiber to collimate the beam. A narrow bandpass filter is placed in the path of the collimated beam to block the Raman signal and to transmit only the probe beam. A high pass filter then reflects the probe beam. A second lens focuses the probe beam onto the sample. In the return path, a notch filter is placed behind the high pass filter to block the Rayleigh scattered probe beam while transmitting the frequency shifted Raman signal. The Raman signal is then focused in a multimode fiber for collection. The diameter of the fiber probe is thus of the order of 10 mm to accommodate the free-space collimating lenses and filters.
Multimode fibers present many advantages for light transmission such as a large fiber core and a large numerical aperture, which lead to a high fiber coupling efficiency and high light gathering feature. For their light high collection efficiency, they are already used for image collection in endoscopy. However, light propagation through a multimode fiber produces a speckle pattern and thus a specific spatial excitation/illumination through a MM fiber is a challenge. Indeed, as the optical field is coupled into the fiber, it excites different fibers modes which propagate along the fiber, possibly exchanging energy between them through the mechanisms of intermodal coupling and finally reaching the output fiber surface where they interfere; generating what is seemingly a random speckle pattern. Multimode fibers have a number of modes M given by M=4V2/p2 f, for M>>1 where V=p·f·NA/l. l is the wavelength of the light, NA is the numerical aperture of the multimode fiber and f is the fiber core diameter. By way of example, a multimode fiber with core diameter 200 mm, NA=0.42 and wavelength 532 nm possess 100,000 modes. A single mode fiber has only 1 mode (M=1).
In addition to this modal spatial scrambling, dispersion effects are also very important in multimode fibers resulting in a temporal spreading of an input light pulse. These two aspects have always limited their use for light transmission, and particularly for fiber-based imaging.
Light scrambling occurring in multimode fibers produce speckle patterns that are similar to light patterns created as a result of propagation in a diffuse medium. There are major differences between propagation in a diffuse medium and in a multimode fiber. One of them, is the forward-only propagation in multimode fiber whereas backward scattering occurs in a diffuse (turbid) medium. Digital phase conjugation methods have been shown to suppress turbidity in a turbid medium: U.S. Pat. No. 5,378,888 (HOLOGRAPHIC SYSTEM FOR INTERACTIVE TARGET ACQUISITION AND TRACKING), US patent application 2011/0122416 A1 (TURBIDITY SUPPRESSION BY OPTICAL PHASE CONJUGATION USING A SPATIAL LIGHT MODULATOR) and publication by C.-L. Hsieh, Y. Pu, R. Grange, and D. Psaltis, “Digital phase conjugation of second harmonic radiation emitted by nanoparticles in turbid media,” Opt. Express 18, 12283-12290, 2010.
A digital phase conjugation technique has been applied to a bundle of near single mode fibers to coherently combine the output of 3 fibers: C. Bellanger, A. Brignon, J. Colineau, and J. P. Huignard, “Coherent fiber combining by digital holography,” Opt. Lett., OL 33, 2937, 2008. The same technique was applied to a fiber with a low number of modes (M=4): M. Paurisse, M. Hanna, F. Druon, P. Georges, C. Bellanger, A. Brignon, and J. P. Huignard, “Phase and amplitude control of a multimode LMA fiber beam by use of digital holography,” Opt. Express 17, 13000-13008, 2009). However such a technique has not been shown to work with very high number of modes. There is thus a need to develop a technique for a very high number of modes. This patent describes a method to use digital phase conjugation for a very high number of modes. Digital phase conjugation is a “single shot” technique meaning that only one digital hologram suffices to control the wavefront to achieve a desired output.
Other techniques are iterative, meaning that the wavefront is optimized sequentially to maximize the response of a beacon (e.g. single detector, fluorescent particle) placed at the output of the multimode fiber. Such iterative techniques applied to multimode fibers are described in Di Leonardo et al. (R. Di Leonardo and S. Bianchi, “Hologram transmission through multi-mode optical fibers,” Opt. Express 19, 247-254, 2011) and T. {hacek over (C)}i{hacek over (z)}már and K. Dholakia, “Shaping the light transmission through a multimode optical fibre: complex transformation analysis and applications in biophotonics,” Opt. Express 19, 18871, 2011.
Yet another method of sending a desired pattern through a turbid medium or multimode fiber is to calibrate the medium i.e. determining the transmission matrix T such that an input image x is transformed into an output image y by the linear relation y=T*x. S. M Popoff, G. Lerosey, R. Carminati, M. Fink, A. C Boccara and S. Gigan et al. “Measuring the Transmission Matrix in Optics: An Approach to the Study and Control of Light Propagation in Disordered Media”, PRL 104, 2010 describes such a method to measure the transmission matrix T.